Intra-surgical optical coherence tomographic imaging of cataract procedures

ABSTRACT

A cataract surgical system includes a laser source to generate a first set of laser pulses; a guiding optic to guide the first set of laser pulses to a target region in an eye; a laser controller to generate an electronic representation of a target scan pattern, and to control the guiding optic to scan the first set of laser pulses according to a portion of the target scan pattern to create a first photo-disrupted region in the target region; and an Optical Coherence Tomographic (OCT) imaging system to generate an image of a portion of the first photo-disrupted region. The laser controller can generate an electronic representation of a modified scan pattern in relation to the image generated by the OCT imaging system, and control the guiding optic to scan a second set of laser pulses according the modified scan pattern.

CROSS REFERENCE TO RELATED APPLICATION

This application is a continuation of, and claims benefit of andpriority to, application Ser. No. 14/715,938, titled “Intra-SurgicalOptical Coherence Tomographic Imaging Of Cataract Procedures,” filed May19, 2015, now pending; which is a continuation of, and claims thebenefit of and priority to, application Ser. No. 13/329,529, titled“Intra-surgical Optical Coherence Tomographic Imaging of CataractProcedures,” filed Dec. 19, 2011, now U.S. Pat. No. 9,066,784. Each ofthe above-identified patent applications is hereby incorporated byreference in its entirety.

TECHNICAL FIELD

This patent document relates to applying Optical Coherence Tomographicimaging systems during cataract procedures. In more detail, this patentdocument relates to intra-surgical Spectral Domain Optical CoherenceTomographic imaging of cataract procedures.

BACKGROUND

Cataract surgery is experiencing a revolution. The precision and speedof many aspects of the cataract procedures has improved dramatically inthe last few years. Pulsed surgical laser systems with femtosecond laserpulses provide very precisely controlled cutting functionalities.Advanced optical systems provide unprecedented control over theplacement and targeting of the laser pulses. In addition, imagingsystems provide high quality images to assist the surgeons to plan andexecute the cataract surgical procedures. However, there is still lot ofroom for the improvement of the cataract surgical systems, especially inthe area of imaging.

SUMMARY

One area where substantial improvement is possible and called for isproviding more extensive and actionable imaging information for thecataract surgeon. The most advanced of today's systems include anOptical Coherence Tomographic (OCT) imaging system. Prior to thecataract surgery, these systems can generate and display an in-depthcurvilinear or cross sectional reference image of the anterior segmentof the eye that includes the cornea, the anterior chamber and the lens.The surgeon can plan the surgical procedure by placing marks on thedisplayed reference image to input characteristic points or end-pointsof the various cuts and regions to be photo-disrupted or photo-treated.An interactive interface of a laser controller can sense these marks andtranslate them into electronic control signals to guide the surgicallaser beam to form the corresponding cuts.

To place the invention in context, it is recalled that a cataractprocedure with a surgical laser system can include the following steps.(1) First, the lens can be photo-disrupted inside the capsular bag byscanning the surgical laser beam according to a photo-disruptionpattern. Depending on the hardness of the cataract, the disruptionpattern, the degree of disruption and the desired surgical outcomes,this process can be referred to as a chop, a fragmentation, or a lysis.(2) Second, a capsular lid or cap can be cut in the capsular bag oranterior capsular layer by a circular capsulotomy, anterior capsulotomyor continuous curvilinear capsulotomy. The capsular lid or cap is formedso that when it is removed, a hole is opened up in the capsular bagthrough which the surgeon can extract or aspirate the photo-disruptedlens from the capsular bag. (3) An access cut can be formed next in thesclera, limbus, or peripheral corneal region of the eye. Through thisaccess cut surgical devices, such as a forceps or a phaco-tip can beinserted into the eye. (4) Next, the capsular lid or cap can be removedby one of the inserted surgical devices to form the aforementionedcapsular opening. (5) Often, the surgical laser does not disrupt thelens completely. In such cases, a phaco-tip can be inserted into thecapsular bag to complete the disruption of the lens by applyingultrasound and chopping. (6) Subsequently, the fragments of the lens canbe removed or aspirated through the capsular opening. (7) Finally, anintra ocular lens (IOL) can be inserted to restore vision in the eye.The order of steps (1) and (2) can be reversed in some procedures.

Cataract procedures can be complemented by forming additional cuts, suchas limbal relaxing incisions or arcuate incisions in the cornea, andvarious additional access cuts.

However, once the photo-disruption by the surgical laser beam started toform the planned cuts, today's imaging systems do not generateadditional images that could provide actionable information or feedbackfor the cataract surgeon. This is primarily due to the fact that withthe modern surgical systems the cataract surgical procedures can last arather short time, such as 10-20 seconds for a capsulotomy or 30-40seconds for a lens photo-disruption. Existing imaging systems are unableto image the photo-disrupted region with sufficient resolution in such ashort surgical time. Even less are they capable of analyzing the imageof the photo-disrupted region to provide actionable feedback, or ofactively modifying the ongoing surgical procedure. Such functionalitieswould require faster imaging performance and additional or differentelectronic and imaging systems.

While imaging and analyzing the affected regions during the shortsurgical times is hard, a feedback based on such an imaging or analysiswould be highly desirable both to improve the precision of the surgicalprocedure and to manage unexpected surgical complications. For example,a capsulotomy may not cut through the entire capsular bag at someportions of the circular cutting line so that the circular lid or capremains attached to the rest of the capsule at these “tags”. Whensubsequently the surgeon attempts to remove the circular lid with aforceps, the capsular bag can be torn at the tags, leading to jaggededges or substantial tears. Had the surgeon been provided with an imageof the incomplete capsulotomy in time, he may have opted for re-scanningthe capsulotomy circle with the laser to cut through the tags ratherthan commencing the removal of the partially undetached lid.

In other cases, when the photo-disruption of the lens is performed, thesurgical laser may be scanned too close to the posterior capsular layer,possibly puncturing it. Such a puncture may necessitate a complexemergency vitrectomy procedure, substantially elevating the risk of theentire cataract procedure. Again, had the surgeon been provided with animaging feedback in a timely manner, she could have modified thescanning pattern to guide the surgical laser beam away from theposterior capsular layer, preventing the vitrectomy.

In yet other cases, the surgical laser system may be mis-calibrated: thelaser controller may have miscalculated the location of the laser pulsesfor a variety of reasons, including optical aberrations, manufacturingtolerance problems of the laser, a mischaracterization of the refractiveproperties of the lens, a pre-operative diagnostic error, a movement orshape-change of the eye, and thermal creep of the components. In anexample, while the surgeon may have placed the marks on a referenceimage to form a surgical cut e.g. 100 microns from the posteriorcapsular layer, the guiding optic may have directed the surgical laserpulses to a location only 50 microns from the posterior capsular layerbecause of the miscalibration, elevating the risk and reducing theprecision and safety of the cataract procedure. As above, had thesurgeon been provided with an image of the progress of the surgery, shecould have discovered the miscalibration before the photo-disruption gotto the dangerously close 50 micron distance from the posterior capsularlayer.

In yet another example, the miscalibration can be caused by the entirelens having moved along the optical axis because of a difference of thepressure in the anterior chamber and in the posterior chamber, orvitreous, located behind the lens along the optical beam path. Thepressure can change for a variety of reasons after the reference imagehas been taken, such as because of the pressure exerted by the patientinterface. Also, the eye being a dynamical system, the internalpressures in the anterior and posterior chambers can change in time,e.g. as the internal pressures equilibrate with the external pressuresover an extended period such as seconds or tens of seconds after thepatient interface is docked to the eye.

In yet another example, the miscalibration can be caused by the lenscurvature having changed because of accommodation. The accommodation canbe caused by the patient before and during the procedure. Surgeonstypically administer drugs to arrest or suppress accommodation, ineffect expanding the pupil. However, these drugs have different effectson different patients and even these different effects follow differenttimelines. Again, in the last two examples, as before, had the surgeonbeen provided by updated or timely images during the procedure, shecould have recognized the miscalibration and could have taken preventiveor corrective action.

A common feature of these and many other possible surgical complicationsis that they become detectable only after the photo-disruption of thetarget tissue has been started. As described above, however, forming animage in an imaging time shorter than a surgical time of 10, 20, 40 or60 seconds can be a great challenge for today's imaging systems,especially if a high resolution image is necessary to provide actionableimaging information. And it can be prohibitively challenging for theexisting systems in addition to analyze the image in order to display afeedback or a suggested corrective action, or in order to activelymodify the scanning of the surgical photo-disrupting laser beam.

The present patent document describes embodiments of cataract surgicalsystems with advanced imaging systems that are configured to image andin some implementations to analyze the regions photo-disrupted by thesurgical laser beam in an imaging time shorter than typical surgicaltimes. These embodiments therefore enable the modification of thecataract surgical procedure in real time, either by the surgeon or bythe surgical system itself, promising a qualitative improvement of theefficacy and safety of modern cataract surgery.

In particular, in an embodiment a cataract surgical system may include alaser source, configured to generate a first set of laser pulses; aguiding optic, coupled to the laser source, configured to guide thefirst set of laser pulses to a cataract target region in an eye; a lasercontroller, configured to generate an electronic representation of atarget scan pattern and to control the guiding optic to scan the firstset of laser pulses according to a portion of the target scan pattern tocreate a first photo-disrupted region in the cataract target region; anda Spectral Domain Optical Coherence Tomographic (SD-OCT) imaging system,configured to generate an image of a portion of the firstphoto-disrupted region; wherein the laser controller is configured togenerate an electronic representation of a modified scan pattern inrelation to the image generated by the SD-OCT imaging system, and tocontrol the guiding optic to scan a second set of laser pulses accordingthe modified scan pattern to create a second photo-disrupted region. Insome embodiments, the imaging system can be a Swept-Source OpticalCoherence Tomographic (SS-OCT) imaging system.

In some embodiments, a cataract surgical system can include a surgicallaser system, configured to generate a surgical laser beam and to guidethe generated surgical laser beam into a cataract target region; a lasercontroller, configured to scan the surgical laser beam in the cataracttarget region to create a photo-disrupted region; a Spectral DomainOptical Coherence Tomographic (SD-OCT) imaging system, configured togenerate an image of the photo-disrupted region for a system operator ina surgical time after the scanning of the surgical laser beam started;wherein the laser controller is configured to stop or suspend thescanning of the surgical laser beam in response to receiving astop-control signal from the system operator in response to thegenerated image.

In some embodiments, an ophthalmic surgical method can includegenerating an electronic representation of a target scan pattern for alens of an eye by a laser controller; generating and scanning a laserbeam in the lens of the eye according to the target scan pattern by asurgical laser system, creating a cut in the lens; generating an imageof a portion of the eye and the cut with a Spectral Domain OpticalCoherence Tomographic imaging system after the scanning of the laserbeam started; generating an electronic representation of a modified scanpattern by the laser controller in relation to the generated image; andgenerating and scanning the laser beam in the lens of the eye accordingto the modified scan pattern by the surgical laser system, creating amodified cut.

In some embodiments, a method of cataract surgery can includecontrolling a scanning of a laser beam in a lens of an eye by a lasercontroller; generating images of a portion of the lens at a rate of atleast 5 frames per second by a Spectral Domain Optical CoherenceTomographic imaging system; and modifying the scanning of the laser beamby the laser controller in response to an analysis of the generatedimages.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A illustrates an embodiment of a cataract laser surgical system.

FIG. 1B illustrates an imaging aided photo-disruption in a lens targetregion.

FIGS. 1C-D illustrate embodiments of a cataract laser surgical system.

FIGS. 2A-C illustrate forming a first and a modified secondphoto-disrupted region.

FIGS. 3A-E illustrate the modification of scan patterns after a surgicalbyproduct has been detected.

FIGS. 4A-B illustrate the modification of the scan pattern after asurgical byproduct has been detected.

FIGS. 4C-D illustrate the modification of the capsulotomy scan patternafter the lens capsule expanded.

FIGS. 5A-D illustrate a re-scanning of an inefficient capsulotomy.

FIG. 6 illustrates an embodiment of a Spectral Domain Optical CoherenceTomographic imaging system.

FIG. 7 illustrates an embodiment of a Spectral Domain Optical CoherenceTomographic imaging system with a dedicated Input-Output board.

FIG. 8 illustrates a dedicated Input-Output board.

FIG. 9 illustrates an embodiment of a Swept Source Optical CoherenceTomographic imaging system.

FIG. 10 illustrates an imaging-aided cataract surgical method.

FIG. 11 illustrates an imaging-aided cataract surgical method.

DETAILED DESCRIPTION

Implementations and embodiments in this patent document describecataract surgical systems that generate timely imaging feedback eitherto assist a surgeon to adjust the surgical procedure based on thefeedback, or to determine and carry out such an adjustment bythemselves.

FIG. 1A illustrates a cataract surgical system 100, including a lasersource 110 to generate a laser beam of a first set of laser pulses.These laser pulses can have a duration or pulse length in the range of1-1,000 femtoseconds or 1-1,000 picoseconds. The energy and power of thelaser beam can be selected to achieve a well controlled photo-disruptionin the selected target region efficiently without causing damage inother ophthalmic tissues such as in the photosensitive retina. Thecataract surgical system 100 can also include a guiding optic 120,coupled to the laser source 110 to guide the first set of laser pulsesthrough a cornea 3 and an anterior chamber 4 to a lens 5 of an eye 1.The lens 5 is enveloped by a capsular layer or bag 6. The guiding optic120 can be configured to guide the laser beam into a cataract targetregion through an objective 122 and a docking unit or patient interface(PI) 124 that can dock onto the eye 1 and immobilize it relative to thecataract surgical system 100 by applying vacuum suction. In someembodiments, there may not be a direct or immobilizing connectionbetween the objective 122 and the eye 1. These embodiments may employinstead eye trackers to correlate the imaging process with possiblemovements of the eye, for example.

The cataract target region can be located in an anterior segment of theeye 1 that includes the cornea 3, the anterior chamber 4, the lens 5 andthe capsular bag 6 of the eye 1. The cataract target region can be, forexample, a circle, cylinder, or slanted cylinder in an anterior capsularlayer or capsular bag of the lens 5 when a capsulotomy is performed. Thecataract target region can also be a large volume fraction of the lens 5to achieve a photo-disruption, a chop or a lysis of the lens 5 or atleast of its nucleus. The cataract target region can also be in thecornea 3, such as an access cut to create a port for the insertion ofcataract surgical devices. In more comprehensive cataract procedures,such as in refractive cataract surgery, additional limbal relaxing cutsor incisions (LRI) or arcuate incisions can be formed as well.

The capsulotomy can have a diameter in the 3-6 mm range, as dictated bythe design of the intra ocular lens, or IOL, to be inserted into thepreserved lens capsule at a z-depth in the 2-4 mm range, where thez-depth is measured along an optical axis of the cataract surgicalsystem 100, using a contact surface of the cornea 3 and the PI 124 as azero reference level for the z-depth. The target region of a lensphoto-disruption can extend from 2-4 mm z-depth to 7-10 mm z-depth, witha diameter of 4-8 mm. Finally, the corneal LRI, arcuate and access cutscan be formed in the 0-2 mm z-depth at the large diameter of 6-14 mm tominimize or altogether avoid directly impacting the field of vision.These numerical ranges show that the challenges of cataract proceduressubstantially exceed those of the purely corneal procedures—such asLASIK—or retinal procedures. Both corneal/LASIK and retinal proceduresare performed in a much narrower z-depth range and in a much smalleroverall surgical volume than cataract procedures. Corneal procedures aretypically restricted to 0.1-0.5 mm z-depth ranges as the thickness ofthe cornea rarely exceeds 1 mm and the photo-disruption typically doesnot cut through the entire cornea to keep the anterior chamber intact.Typical diameters of corneal cuts can be in the 2-3 mm range. Whileretinal procedures are performed deep in the posterior segment of theeye 1 at a large z-depth, the range of z-depths where the cuts areformed is typically less than 2 mm, the overall thickness of the retinallayers of interest.

In contrast, cataract procedures typically involve the photo-disruptionin most or all of the above described cataract target regions, both inthe cornea 3 and in the lens 5. Therefore, cataract procedures caninvolve cuts in a z-depth range of 4 mm or larger, sometimes 6 mm orlarger. These cataract z-depth ranges are substantially larger than theabove described z-depth ranges of the corneal or retinal procedures.Further, the diameter of the cataract-related cuts also exceeds that ofthe corneal cuts substantially. Therefore, forming cataract cuts posessubstantially harder challenges for the design of a cataract surgicalsystem, including its imaging system, than forming corneal cuts posesfor the design of a LASIK system, or forming retinal cuts poses for thedesign of a retinal surgical system.

The cataract surgical system 100 can also include a laser controller 130to generate an electronic representation of a target scan pattern and tocontrol the guiding optic 120 to scan the first set of laser pulsesaccording to a portion of the target scan pattern to create a firstphoto-disrupted region in the cataract target region.

As described above, the cataract target region can be a proximity of ananterior capsular layer and the target scan pattern can be a set oftarget points on a circle or cylinder in the anterior capsular layer ofthe lens 5 for a circular capsulotomy, an anterior capsulotomy, or acurvilinear capsulotomy.

Or, the cataract target region can be a portion of the lens 5 and thetarget scan pattern can be a set of target points on radial chop planes,cylinders, a spiral pattern, or a mesh pattern to inducephoto-disruption of the lens 5 itself. The points of the target scanpattern can be defined e.g. by their radial or (x,y,z) coordinates.These coordinates can be electronically represented in a processor,based on executables stored in a corresponding memory of the lasercontroller 130.

The cataract surgical system can also include a Spectral Domain OpticalCoherence Tomographic (SD-OCT) imaging system 200 to generate an imageof a portion of the first photo-disrupted region, created by thescanning of the surgical laser beam. The SD-OCT imaging system 200 canbe configured to couple an imaging beam into the guiding optic 120 to beguided into the eye 1 and to receive a returned imaging beam from theguiding optic 120. The SD-OCT imaging system 200 can be configured togenerate the image or images of the first photo-disrupted region duringthe surgery to provide timely or actionable feedback for the surgeon orfor the laser controller 130, as described below in detail.

FIG. 1B illustrates an operation of the cataract surgical system 100. Inthis example, the laser controller 130 can generate the electronicrepresentation of a target scan pattern 302 in the cataract targetregion, an arc close to the posterior capsular layer. The guiding optic120 can focus and scan the first set of laser pulses of a surgical laserbeam 304, generated by the laser source 110, through the points of thetarget scan pattern 302 to create a first photo-disrupted region 306.The first photo-disrupted region 306 in this example can consist of aset of bubbles or cavitation bubbles, formed at the points of the targetscan pattern 302. After the photo-disruption started, the SD-OCT imagingsystem 200 can scan an imaging beam 308 through the cataract targetregion to generate an image of the first photo-disrupted region 306. Insome implementations, the imaging beam 308 and the surgical laser beam304 can be scanned or guided by the same shared guiding optic 120. Inother implementations, only part of the optical pathway can be sharedand the imaging beam 308 can be partly scanned by an additionalnon-shared imaging-guiding optic. All these designs are embodiments ofthe guiding optic 120.

If the image generated by the SD-OCT imaging system 200 indicates thatthe procedure is progressing as planned, such as the photo-disruptedbubbles 306 are formed according to the target scan pattern 302 andwithout unintended consequences, the laser controller 130 can continuescanning the surgical laser beam 304 along the original target scanpattern 302. However, if the image indicates that there is a deviationfrom the planned procedure, the laser controller 130 can respond bygenerating an electronic representation of a modified scan pattern andcontrol the guiding optic 120 to scan a second set of laser pulsesaccording the modified scan pattern to create a second photo-disruptedregion, as will be illustrated in subsequent figures.

In some embodiments, there may be no direct coupling between the SD-OCTimaging system 200 and the laser controller 130. In these embodiments,the SD-OCT imaging system 200 can display the image of the firstphoto-disrupted region 306, and a system operator, such as the surgeoncan enter modified scan parameters to cause the laser controller 130 togenerate the electronic representation of the modified scan pattern.

FIG. 1A illustrates that in some embodiments, the SD-OCT imaging system200 can include an OCT image processor 201 that can analyze thegenerated image after the photo-disruption started. In some embodiments,the OCT image processor 201 can display the results of the analysis forthe surgeon to provide a timely and actionable feedback during thesurgery, so that the surgeon can enter a modified scan pattern into thelaser controller 130. In the example of FIG. 1B, the OCT image processor201 may be configured to measure the distance of the firstphoto-disrupted region 306 from the posterior capsular layer and if thedistance becomes smaller than a preset margin of safety, then displaywarning signals for the surgeon.

In some embodiments, the SD-OCT imaging system 200 can be coupled to thelaser controller 130, as in FIG. 1A, or the OCT image processor 201 canbe a self-standing unit, directly coupled to SD-OCT imaging system 200and to the laser controller 130, as in FIG. 1C. In these embodiments,the OCT image processor 201 can generate control signals in relation tothe image of the first photo-disrupted region 306 and can apply thegenerated control signals to the laser controller 130 to cause the lasercontroller 130 to generate the electronic representation of the modifiedscan pattern. The OCT image processor 201 can be fully or partiallyintegrated into the SD-OCT imaging system 200 as in FIG. 1A.

FIG. 1D illustrates that in some embodiments the OCT image processor 201can also be overlapping or even integrated with the laser controller130. The embodiments of FIGS. 1A, 1C and 1D illustrate that thesoftware-related functions of processing the OCT image and generatingthe modified scan pattern can be partially or fully executed by amulti-purpose processor that can be housed in either the SD-OCT imagingsystem 200, or the laser controller 130, or in a block integrating both,or can be a self-standing block, separate from both.

As mentioned above, the unusually large z-depth range of cataractsurgeries that can be larger than 4 mm or in some embodiments largerthan 6 mm may require the use of a substantially more complex SD-OCTimaging system 200 than what is employed in corneal or retinal systems.Accordingly, in some embodiments the SD-OCT imaging system 200 can beconfigured to have an imaging or z-depth range Lmax larger than 4 mm,such as in the range of 4-20 mm. In other embodiments, the imaging orz-depth range Lmax can be larger than 6 mm, such as in the range of 6-10mm.

Lmax, the imaging or z-depth range of SD-OCT imaging system 200 candepend on the wavelength λ of the imaging laser beam 308, the wavelengthresolution δλ, the Nyquist frequency Nf, the focal length f and thepupil d of the SD-OCT-imaging system 200, as described below in detail.Accordingly, embodiments of the SD-OCT imaging system 200 can bedesigned with parameters λ, δλ, Nf, f and d such that the imaging orz-depth range is larger than 4 mm, or in some embodiments larger than 6mm.

The difficulty of extending the imaging depth range of a system from 1-2mm to 4 mm or more can be also appreciated from the fact that someexisting systems that require larger than 2 mm imaging ranges achievethis not by involving more advanced optics as that would have beenprohibitively hard. Instead, these systems employ a conventional imagingsystem with a less than 2 mm imaging range and boost the imaging rangeof this conventional system by taking several images at adjacentz-depths separated by about 2 mm and generate a single image with thelarger range by integrating the adjacent-depth images using a compleximage recognition and processing circuitry. Such systems can beelectronically complex and the image integration slows down the speed oftheir performance considerably. To avoid the substantial slowing down ofthe imaging performance and the need for complex electronics,implementations of the SD-OCT imaging system 200 achieve the imagingdepth range of more than 4 or 6 mm without integrating two or moreimages.

For clarity, it is noted that it is customary to distinguish between twotypes of imaging scans: A-scans and B-scans. An A-scan refers to animage of the target in a range of z-depths corresponding to a singletransverse (x,y) coordinate, in the reference frame whose z axis isaligned with an optical axis of the guiding optic 120. An A-scan can beobtained by directing an imaging beam of an imaging system to a single(x,y) point of the target and collecting the imaging informationcorresponding to different z-depths.

Some imaging systems generate an A-scan by indeed scanning the z imagingdepth-range and recording the image data for different z-depthssequentially. However, while SD-OCT imaging systems, as explained below,collect the image data for different z-depths simultaneously, i.e.without scanning in the z direction, yet their images are still oftenreferred to as A-scans.

A B-scan refers to a set of A-scans that correspond to a set or line of(x,y) points, collected as the imaging beam is scanned along atransverse line or in a transverse scanning pattern. A typical B-scanwith regular (x,y) resolution can include 500-2,000 A-scans. A B-scanwith high (x,y) resolution can include 1,000-3,000 A-scans. Particularlyhigh (x,y) resolution B-scans can include 2,000-5,000 or 2,000-16,000A-scans. Typically, the B-scan can include these A-scans integrated intoa cross sectional, circular or cylindrical image of the target. As such,a B-scan can provide a substantially more detailed and thussubstantially more actionable feedback imaging information for thesurgeon than an individual A-scan. Accordingly, in the embodiments ofthe cataract surgical system 100 an image of the first photo-disruptedregion and the second photo-disrupted region can refer to a B-scan thatcan include 500-2,000, 1,000-3,000, 2,000-5,000, or 2,000-16,000A-scans.

OCT imaging systems can be categorized into two classes: Time Domain, orTD-OCT imaging systems, and Spectral Domain, or SD-OCT imaging systems.TD-OCT imaging systems use an imaging light beam with a bandwidthsuitable to define short pulse lengths and gather the imaginginformation from different z-depths sequentially, in essence scanningalong the z axis. In contrast, SD-OCT imaging systems use an imaginglight beam with a bandwidth where the different wavelength spectralcomponents capture and carry imaging information representing differentz-depth in parallel, at the same time. This allows the SD-OCT imagingsystems to gather the imaging information from different z-depthssimultaneously, in parallel. Parallel sensing of the z-depth imaginginformation accelerates the performance of the SD-OCT imaging systems bya factor of 10-1,000 relative to the TD-OCT imaging systems. This fasterperformance of the SD-OCT imaging systems can be utilized in severalembodiments, as described next.

In terms of imaging times, this accelerated performance translates toembodiments of the SD-OCT imaging system 200 being able to generate aB-scan image after the photo-disruption started in an imaging time lessthan a surgical time. The imaging time can be less than 1 second, suchas in the range of 0.1 msec-1 sec. In some embodiments the imaging timecan be less than 0.1 second, such as in the range of 1 msec-0.1 sec.These short imaging times mean that the SD-OCT imaging system 200 cangenerate images that can provide timely and thus useful feedback aboutthe progress of the cataract procedure for the surgeon so that thesurgeon can modify the surgical procedure in response to the feedback.This modification can include entering a modified target scan pattern.

The next level of utility is offered by some embodiments of the SD-OCTimaging system 200 that can provide feedback images not only once butrepeatedly during the cataract surgery. Such systems can providevaluable timely feedback regarding the development, location and growthof the first photo-disrupted region 306, thus offering qualitativeimprovement in the precision, performance and safety of the cataractsurgical system 100.

Some embodiments of the SD-OCT imaging system 200 can offer furtherqualitative improvements. They can provide not only a few updated imagesduring the cataract surgery, but an essentially live image of theprogress of the procedure. An essentially live feedback can deliverhighly valuable, timely, and actionable information for the surgeon tomonitor the progress of the surgery, improve the surgical precision,detect undesirable outcomes early and react to them in real time.

An often used refresh rate of live video images is about 24frames/second. Therefore, imaging systems that can provide images at arefresh rate or frame rate of 20 to 25 frames/second or higher canprovide images that will appear essentially live for the surgeon.Whereas systems with a frame rate or refresh rate considerably less than20-25 frames/second may not be perceived as live video imaging, butrather as jerky, jumpy images, possibly even distracting the surgeonfrom the cataract procedure.

In this context, because TD-OCT imaging systems acquire z-depth imaginginformation sequentially, they may be able to generate only lowresolution B-scans with a refresh rate of only one or few frames persecond. TD-OCT imaging systems that are expected to provide images withhigher resolution may be forced to scan and refresh the images at aneven lower rate, well below one frame/second. Such distinctlyslower-than-live feedback images appear jerky for the surgeon and caneven be a distraction. Moreover, the slow scanning speed and resultingslow refresh rate can make some TD-OCT imaging systems to displayartifacts, such as steps or discontinuous jumps in the image of a layerthat in reality is smooth.

In contrast, SD-OCT systems gather image data at an (x,y) point from allz-depths simultaneously, in parallel. These images are sometimes stillcalled A-scans, even though no sequential z-scanning is involved.Because of the parallel or simultaneous nature of gathering theimage-data from different depths, embodiments of the SD-OCT system 200can acquire the A-scans 10-1,000 times faster than TD-OCT imagingsystems, as discussed above. In particular, quality SD-OCT imagingsystems 200 can acquire 10,000-100,000 A-scans per second, orequivalently, can have an A-scan acquisition rate of 10-100 kHz. Highquality SQ-OCT imaging systems 200 can have an A-scan acquisition rateof 30-300 kHz, and particularly high quality SD-OCT imaging systems 200can have an A-scan acquisition rate of 100 kHz-1,000 kHz, much exceedingthe A-scan acquisition rate that can be achieved by TD-OCT imagingsystems.

Clearly, the A-scan acquisition rate, or number of A-scans/sec, isapproximately equal to the number of A-scans/B-scan times the number ofimages/sec, the image refresh rate. For example, at the quality A-scanacquisition rate of 10,000-100,000 A-scan/sec, or 10-100 kHz, imageswith the regular (x,y) resolution of 500-2,000 A-scan/B-scan can becaptured at image refresh rates in the range of 5-200 frames/sec thatincludes the refresh rate range of 20-200 frames/sec. In anotherexample, at the high quality A-scan acquisition rate of 30-300 kHz,images with the high (x,y) resolution of 1,000-3,000 A-scan/B-scan canbe captured at image refresh rates in the range of 10-300 frames/secthat includes the 25-300 frames/sec range. Finally, at the particularlyhigh quality A-scan acquisition range of 100-1,000 kHz, images with theparticularly high (x,y) resolution of 2,000-5,000 A-scans/B-scan can begenerated with image refresh rates in the range of 25-500 frames/sec.

These examples show that embodiments of the SD-OCT imaging systems 200with various pairings of the A-scan acquisition rate quality and theA-scan/B-scan resolutions ratios can provide image refresh rates thatare distinctly above the 20 frames/sec threshold live video rate andthus offer qualitative advantages over the TD-OCT imaging systems.

It is noted that embodiments of the SD-OCT imaging system 200 can beoperated at image refresh rates lower than the live video rate,typically when operated with a particularly high resolution and aregular A-scan acquisition rate quality. Such embodiments may be usedwhen the operator of the SD-OCT imaging system 200 calls for theparticularly high resolution setting for medical reasons, willinglyforegoing the live-video-rate capability of the SD-OCT imaging system200.

The overall amount of image data can be captured in other ways as well.The specific design parameters of the SD-OCT imaging system 200, such asthe full length of its sensor array control the z-directional distanceof the resolved points, the z-resolution. This z-resolution can be, forexample, a 5 micron z-directional distance between data points,translating to 1,000 z-depth points in a typical Lmax=5 mm z-depthrange. In a regular (x,y) resolution embodiment, where a B-scan contains500-2,000 A-scans, often spaced apart also by about 5 microns in the(x,y) plane, this embodiment can generate an image with animage-resolution of 0.5-2 million pixels per image. Other embodimentsmay be able to capture images with an image resolution of 1-3 million,2-5 million, or even 2-10 million image points/frame, still providingthe images at a live video rate of 20-200, 25-300, or 25-500 frames/secor faster.

Because of these remarkably high (x,y), z- and image resolutions,embodiments of the SD-OCT imaging systems 200 can capture and displaycomplex, sharp and detail-rich images. For example, the B-scans caninclude scanning the lens 5 along several circles, radial rays, spirals,and two dimensional (2D) transverse or lateral scanning grids in the(x,y) plane. These detailed images allow the SD-OCT imaging system 200to map out the actual shape of the lens 5 instead of using models andrelying on assumptions about its geometry and shape.

It is noted here that displaying the OCT images also takes time. Thus,the refresh rate of the image display, limited by the speed of theelectronic performance of a display unit of the SD-OCT imaging system200, might be lower than the rate of the OCT image acquisition-unit. Inthis context, the above cited refresh rates characterize the speed ofimage-acquisition by the SD-OCT imaging system 200, not the display rateof the display unit that can be slower, depending on the electronic anddata-transfer limiting factors.

Given that the imaging speeds of the SD-OCT and TD-OCT imaging systemsare on opposite sides of the live video-rate of 20-25 frames/second,embodiments of the cataract surgical system that include the SD-OCTimaging system 200 can be capable of providing timely and smooth livefeedback information for the surgeon that are free of motionalartifacts, whereas those that use typical TD-OCT imaging systems are notcapable of providing such smooth live feedback for the surgeon and areprone to display motional artifacts.

A final factor, impacting the long term performance of embodiments ofthe SD-OCT imaging system 200 is that SD-OCT imaging systems do not havemoving parts and thus their reliability and serviceability is quitesatisfactory. In contrast, TD-OCT systems have rapidly moving parts,associated with the movement of a reference mirror in a reference arm ofthe TD-OCT apparatus. Obviously, the presence of moving parts in theTD-OCT systems increases the chance of malfunction and misalignment,thus possibly decreasing their overall performance, demanding morefrequent field-service and still facing the higher likelihood oflong-term performance degradation.

In sum, SD-OCT imaging systems are qualitatively different from TD-OCTsystems, especially for cataract applications, at least for thefollowing reasons. (1) SD-OCT imaging systems are configured to providelive imaging or feedback images at refresh rates of 20-200, 20-300, or20-500 frames/sec, useful for high precision cataract surgicalprocesses, whereas TD-OCT systems are not. (2) SD-OCT imaging systemscan provide images at live video rates with high (x,y) resolution of500-2,000, 1,000-3,000, or 2,000-5,000 A-scan/B-scans or higher, whereasTD-OCT imaging systems cannot. (3) SD-OCT imaging systems can beoperated with a quality A-scan acquisition rate of 10-100 kHz, 30-300kHz, or 100-1,000 kHz, whereas TD-OCT system may not. (4) SD-OCT imagingsystems are well suited to provide detail-rich images with high imageresolution at live video rates, such as with 0.5-2, 1-3 or 2-5 millionimage points, whereas TD-OCT system are not. (5) SD-OCT imaging systemscan provide images so rich in detail that an overall image of the lens 5can be formed without using a model of the lens 5, whereas TD-OCT systemcannot. (6) SD-OCT imaging systems typically do not display motionalartifacts, whereas TD-OCT systems are likely to do so. (7) Finally,SD-OCT imaging systems require only infrequent field service andmaintenance, such as only every 6 or 9 months, whereas TD-OCT imagingsystems with their moving parts typically require field services andmaintenance much more often.

An advantage of the SD-OCT imaging system 200 providing one or morefeedback images or a feedback video of the cataract target region ingeneral and the first photo-disrupted region 306 in particular is thatthe surgeon can react to the feedback and modify the surgery bygenerating a modified scan pattern in response to the provided feedbackimages or video. The modified scan pattern can be generated in differentmanners after the laser controller 130 generated the electronicrepresentation of the original target scan pattern 302 according to aninitial input received from a system operator.

In some embodiments, the laser controller 130 can generate theelectronic representation of the modified target scan pattern accordingto a modification input also received from the system operator. In suchembodiments, the modification input can be generated by the systemoperator in response to analyzing the image of the portion of the firstphoto-disrupted region 306. For example, the surgeon can study the imageof the first photo-disrupted region 306, discover the formation of anundesirable gas bubble that would scatter the surgical laser beamexcessively, and cause the generation of a modified scan pattern thatsteers the surgical laser beam 304 clear from the gas bubble.

In other embodiments, the SD-OCT imaging system 200 can displaycalibration marks to assist the surgeon or system operator to calibratea location of the first photo-disrupted region 306 relative to thetarget scan pattern 302. Such calibration marks can be associated withcharacteristic points of the image, such as the apex of the cornea 3 orthe lens 5.

In yet other embodiments, the SD-OCT imaging system 200 can include theOCT image processor 201. The OCT image processor 201 can analyze theimage of the first photo-disrupted region 306 and display a feedback forthe system operator. For example, if the OCT image processor 201 sensesa miscalibration, i.e. that the first photo-disrupted region 306 isformed at a distance from where the target scan pattern 302 would haverequired, it can send a miscalibration feedback signal to the surgeon,who can then decide to stop the procedure and recalibrate the system, orto enter a modified scan pattern that compensates the miscalibrationdistance.

In yet other embodiments, the SD-OCT imaging system 200 can take areference image before the first set of surgical laser pulses weregenerated and a feedback image after the first set of laser pulsesgenerated the first photo-disrupted region 306. The OCT image processor201 can determine the differences between the two images and displayindications of these differences. The determined and displayeddifferences can allow the surgeon or the OCT image processor 201 tomonitor the progress of the cataract surgery, to track discrepanciesrelative to the target scan pattern 302, including the appearance ofunintended surgical byproducts, such as gas bubbles in a timely manner,and to cause the generation of the modified scan pattern in response tothe image and displayed differences.

Finally, in some embodiments the OCT image processor 201 can play a moreadvanced and active role. For example, the OCT image processor 201 canactively analyze the image of the first photo-disrupted region 306 andgenerate a control signal to cause the laser controller 130 to generatethe electronic representation of the modified scan pattern withoutnecessarily waiting for a modification input from the surgeon. Forexample, if the OCT image processor 201 discovers an imminent high risksituation, such as the first photo-disrupted region getting formed tooclose to the posterior capsular layer, then it can stop the cataractprocedure without waiting for the outcome of a slow and time-consuminginteraction with the surgeon.

The above five embodiments can be combined in various ways. For example,upon determining from the image that a high risk situation is imminent,the OCT image processor 201 can both display a feedback signal for thesurgeon and prepare a control signal for the laser controller 130 tostop the procedure. The OCT image processor 201 can then wait for amodification input from the surgeon for a predetermined time. In theabsence of such a modification input within the predetermined time, theOCT image processor 201 can proceed with an action on its own to preventthe impending high risk situation and send the control signal to thelaser controller 130 to stop the procedure without further waiting foran input from the surgeon.

FIGS. 2A-B illustrate the particular example of the laser controller 130having planned the target scan pattern 302 in the posterior region ofthe lens 5. The cataract surgical system 100 may apply the surgicallaser beam 304 according to the target scan pattern 302, creating thefirst photo-disrupted region 306. However, the subsequent imaging of thefirst photo-disrupted region 306 with the imaging laser beam 308 mayreveal that the first photo-disrupted region 306 may have been formedmisplaced by a deviation 310 from its intended target scan pattern 302.The deviation 310 can be a deviation of a location, an orientation or ashape of the first photo-disrupted region 306 relative to the targetscan pattern 302. This misplacement or miscalibration can happen for avariety of reasons: a mistaken input by the surgeon, a manufacturingimprecision of the production process of the guiding optic 120, athermal expansion of the components, a movement by the patient after theinitial imaging, a deformation of the eye caused by the pressure of thepatient interface 124, or a mischaracterization of the refractiveproperties of the eye 1 during a preoperative diagnostic process, amongothers.

FIG. 2B illustrates that in some embodiments, the SD-OCT imaging system200 can display the feedback image of the target scan pattern 302 andthe image of the first photo-disrupted region 306 without an analysis.From this feedback image the surgeon can visually determine thedeviation 310 and enter a compensating modification input to cause thelaser controller 130 to modified scan pattern 312. In other embodiments,the SD-OCT imaging system 200 can display calibration marks to assistthe surgeon's analysis.

In other embodiments, the OCT image processor 201 can determine thedeviation 310. Then, in some embodiments the OCT image processor 201 candisplay the determined deviation 310 as well as the feedback andrecommendation for the surgeon, such as a suggested magnitude anddirection of a shift of the scan pattern, or a warning sign for thesurgeon. In these embodiments, the surgeon can enter a modificationinput into the laser controller 130 to cause the generation of amodified scan pattern 312 with the goal of reducing the deviation 310 tothe reduced deviation 310 r. In yet other embodiments, the OCT imageprocessor 201 can directly signal the laser controller 130 to generatethe electronic representation of the modified scan pattern 312 to reducethe determined deviation 310 to the reduced deviation 310 r, all withoutwaiting for a modification input from the surgeon.

FIG. 2B illustrates that the generation of the modified scan pattern 312can take into account the miscalibration of the guiding optic 120 in anyof the above embodiments. The modified scan pattern 312 can be shiftedfrom the target scan pattern 302 by about the miscalibration of theguiding optic 120 so that when the surgical laser beam 304 forms asecond photo-disrupted region 314 misplaced from the modified scanpattern 312 because of the miscalibration of the guiding optic 120, thesecond photo-disrupted region 314 ends up close to the originallyintended target scan pattern 302, reducing the deviation 310 to thereduced deviation 310 r.

FIG. 2C illustrates a related embodiment, where the OCT image processor201 can determine whether the first photo-disrupted region 306 gotformed or extended inadvertently into a region of risk 316. The firstphoto-disrupted region 306 being formed in the region of risk 316endangers the integrity of the posterior capsule layer and can breachit, puncturing the capsule 6 and necessitating a complex vitrectomyprocedure. This would substantially elevate the risk of the cataractprocedure.

To preempt such a breach, in various embodiments the OCT image processor201 can analyze the feedback image or images, or the essentially liveimaging of the SD-OCT imaging system 200 to monitor whether the firstphoto-disrupted region 306 was formed too close to or in the posteriorcapsule layer, in the region of risk 316. If the OCT image processor 201senses that the first photo-disrupted region 306 has been formed in thisregion of risk 316, then the OCT image processor 201 can either displaya warning feedback for the surgeon, or can generate a control signal forthe laser controller 130 to cause the generation of the modified scanpattern 312. In all the above embodiments, the laser controller 130 cangenerate the modified scan pattern 312 and direct the surgical laserbeam 304 accordingly to form the second photo-disrupted region 314outside the region of risk 316.

In some embodiments, the OCT image processor 201 and the lasercontroller 130 can be partially or fully integrated. For example, asingle integrated processor can perform both the image processingfunction of the OCT image processor 201 and the scan pattern generatingfunction of the laser controller 130.

FIGS. 3A-E illustrate an embodiment where the target scan pattern 302 isa chop pattern. In some cases the cataract surgeon may choose to chopthe lens into 2, 4 or 6 portions to facilitate their removal oraspiration through the capsulotomy.

FIGS. 3A-B illustrate that in the design phase of the cataract surgery,upon receiving the corresponding input from the surgeon, the lasercontroller 130 may generate a target scan pattern 302 that consists ofpoints on two chop planes, formed in the (x,z) and (y,z) planes, forexample. In an ideal cataract procedure, when the first set of laserpulses are applied to this target scan pattern 302, the generated firstphoto-disrupted region 306 includes four chop planes, chopping the lensinto four segments along these chop planes.

FIG. 3C-D illustrate that, in non-ideal cases, after the first set oflaser pulses of the surgical laser beam 304 have been directed to thepoints of the target scan pattern 302 and the first photo-disruptedregion 306 started to form, an unintended surgical byproduct 320 canform as well. This surgical byproduct 320 can be a group of the freshlyformed photo-disruption bubbles coalescing into a single large bubblethat can scatter or redirect the subsequent laser pulses in unintendeddirections, such as towards the retina, possibly causing damage andphototoxicity. Therefore, the precision of the photo-disruption processcan diminish and its risk can increase if subsequent laser pulses aredirected into the gas bubble 320.

Such as unintended development can be preempted by an OCT imageprocessor 201 that can recognize the formation of the gas bubble oranother surgical byproduct 320 from analyzing the images of the SD-OCTimaging system 200. And since the SD-OCT images can be generated at anessentially live rate, the OCT image processor 201 can relay thisrecognition as a timely and actionable feedback either for the surgeonor for the laser controller 130, prompting a modifying response, asdescribed next.

FIGS. 3C-D illustrate that the feedback can take the form of the OCTimage processor 201 analyzing the image, determining a recommendeddegree of rotation of the chop pattern and displaying the recommendedrotated chop pattern for the surgeon to enter the correspondingmodification input. In other embodiments, the OCT image processor 201can apply a control signal to the laser controller 130 directly togenerate an electronic representation of a rotated chop pattern as themodified scan pattern 312 such that the rotated chop pattern 312 isnon-overlapping with the gas bubble 320. The rotated chop pattern 312can extend through the entire z-depth range of the lens 5, in effectstarting the procedure over, or can be a partial chop pattern,continuing the cutting from the z-depth where the original target scanpattern 302 was abandoned.

FIG. 3E illustrates that the OCT image processor 201 could recommend todirect the surgical laser beam 304 to the points of the modified orrotated scan pattern 312 to form the second or rotated or modifiedphoto-disrupted region 314 that does not overlap with the coalescedbubble 320. This embodiment thus avoids the surgical laser beam 304being scattered into unintended directions, reducing the risk of thecataract procedure.

FIG. 4A illustrates an alternative embodiment of forming a modified ornon-overlapping scan pattern 312 in response to the OCT image processor201 sensing the emergence of the surgical byproduct 320. Here, themodified or non-overlapping scan pattern 312 does not include rotatedchop planes. Instead, the overlap is avoided by the surgical laser beam304 being scanned according to the unchanged target scan pattern 302 butthe laser controller 130 blanking out the laser pulses that would bedirected into the surgical byproduct bubble 320. The laser controller130 can un-blank the laser pulses to hit all the points of the targetscan pattern 302 once the scanning moved past the surgical byproductbubble 320.

FIG. 4B illustrates yet another embodiment where once the OCT imageprocessor 201 recognizes that continuing the scanning of the surgicallaser beam 304 along the original target scan pattern 302 would create afirst photo-disrupted region 306 overlapping with the surgical byproductbubble 320 by analyzing the feedback images or the live rate imagestream of the SD-OCT imaging system 200, it can simply stop the scanningof the surgical laser beam 304. In some embodiments, after the stoppingthe OCT image processor 201 or the laser controller 130 can display aprompt for the system operator, asking for a modification input orcorrective action.

In some surgical scenarios, multiple bubbles 320-i may form more or lesssimultaneously. Therefore, some embodiments of the cataract surgicalsystem 100 can be configured to generate a modified scan pattern 312that avoids multiple bubbles simultaneously.

FIGS. 4C-D illustrate that in some embodiments the OCT image processor201 can be configured to analyze a portion of the image that is distinctfrom the first photo-disrupted region 306, and to generate a feedbackbased on this analysis. In this particular example, the formation of thefirst photo-disrupted region 306 can cause the lens 5 and thus the lenscapsule 6 to expand, as indicated by the arrow. The SD-OCT imagingsystem 200 may image this expanded anterior capsular layer 322. The OCTimage processor 201 may analyze this image and determine the shift ofthe location of the expanded anterior capsular layer 322.

Knowing the location of the anterior capsular layer is important for acataract procedure because in some embodiments the capsulotomy targetscan pattern 324 is placed to cut through the anterior capsular layer.If the capsulotomy target scan pattern 324 is placed according to animage taken before the capsular expansion, then the surgical laser beam304 will attempt to create the capsulotomy at an incorrect location.

FIG. 4D illustrates that to prevent this from happening, the OCT imageprocessor 201 can determine the shift of the location of the expandedanterior capsular layer 322 and either display this shift for the systemoperator to enter a modification input into the laser controller 130, orsignal this shift directly to the laser controller 130, in either caseto cause the laser controller to generate a modified capsulotomy scanpattern 326 that properly generates the capsulotomy in the expandedcapsule.

FIGS. 5A-D illustrate another embodiment in connection to forming alaser-assisted capsulotomy or incision. In this embodiment, the OCTimage processor 201 can analyze the repeated feedback images or livestreaming images and identify an uncut portion or “tag” 330 within thefirst photo-disrupted region 306 where the efficiency of thephoto-disruption was limited and thus the anterior capsular membrane wasnot cut fully through, leaving the uncut portion or “tag” 330 behind.Such tags 330 can elevate the risk of tearing the capsule or capsularmembrane when the capsular lid or cap is removed by the surgeon. Toprevent such an undesirable or high risk outcome, the OCT imageprocessor 201 can either display a recommendation for the surgeon how toremove the tag 330 by scanning along an additional modified scan pattern312, or can apply a control signal to the laser controller 130 directlyto generate the electronic representation of the modified scan pattern312 to rescan at least the tag 330 to generate a re-scanned tag-region332 as the second photo-disrupted region 314. With this re-scanning, thefirst photo-disrupted region 306 and the second photo-disrupted region314 can form a completed photo-disrupted region 334, in this case acompleted capsulotomy 334, allowing the surgeon to remove the capsularlid or cap with a minimal risk of tearing.

FIG. 6 illustrates that in some embodiments the SD-OCT imaging system200 can comprise a Spectrometer-Based-OCT (SB-OCT) imaging system 200that includes a broad-band light source 210 to generate a broad-bandlight or laser beam. The broad-band beam can have a mean wavelength λ₀and a relatively broad bandwidth W_(source). In some typical examples,λ₀ can be in the 800-1100 nm range, and W_(source) can be in the 10-150nm range.

The broad-band beam can be coupled into a beam guidance system 211 thatcan include a 1^(st) beam splitter 212. The 1^(st) beam splitter 212 cansplit the broad-band beam into an image beam 222 and a reference beam224. The image beam 222 can be guided by a 2^(nd) beam splitter 214 intothe main optical pathway of the guiding optic 120, and from there onthrough the objective 122 and possibly the patient interface 124 to theeye 1. The beam guidance system 211 can also guide a returned image beam222 r from the eye 1 to the 1^(st) beam splitter 212. The image beam 222and returned image beam 222 r were previously referred to jointly as theimaging laser beam 308.

In addition, the beam guidance system 211 can guide the reference beam224 to a reference mirror 216, guide a returned reference beam 224 rfrom the reference mirror 216, and combine the returned image beam 222 rand the returned reference beam 224 r into a combined beam 226 at the1^(st) beam splitter 212. The combined beam 226 carries the imaginginformation from the eye 1 in the interference of the returned imagebeam 222 r and the returned reference beam 224 r. Some embodiments mayuse other types of delay elements in place of or in conjunction with thereference mirror 216. Others may use yet another beam splitter forcombining the returned image beam 222 r and returned reference beam 224r. In some embodiments, the beam guidance system 211 can include aMach-Zehnder interferometer. Such systems may have favorable noisereduction properties.

TD-OCT imaging systems capture the imaging data from the differentz-depths in the z imaging range sequentially by moving the referencemirror 216 in a corresponding distance range for each (x,y) pointseparately, essentially like a Michelson-Morley interferometer. Incontrast, SD-OCT imaging systems use the different spectral componentsof the broad-band imaging light to capture the imaging data fromdifferent z-depths in parallel. The SD-OCT imaging systems can bethought of as many Michelson-Morley (MM) interferometers operating atdifferent wavelengths in parallel. Since the MM systems operating atdifferent wavelengths image the eye 1 at different z-depths, thecombined beam 226 of an SD-OCT system 200 carries the imaging data orinformation from all z-depths of the eye 1 simultaneously and thus doesnot require the movement or scanning of any mechanical system component,such as the reference mirror 216. As discussed above, this absence ofmechanical scanning for SD-OCT systems translates to an acceleration ofthe imaging speed by a factor of at least 10, or more typically100-1,000 relative to TD-OCT imaging systems.

The SD-OCT imaging system 200 can also include an OCT camera 230,configured to receive the combined beam 226. To recover the imageinformation for all z-depths, the combined beam 226 can be decomposedinto its spectral components by a spectral decomposer 231 such as aprism or grating. Each spectral component with δλ bandwidth around awavelength λ and the interference information they carry can be sensedin parallel by individual sensors of a sensor array 232, the sensorsbeing separated by a distance d′ from each other. The interferenceinformation sensed by the sensors individually can then be used toreconstruct the image of the entire z-depth range by aFast-Fourier-Transformer (FFT) system 233 to generate a Fouriertransform from the sensed spectral components. In effect, theinterference data or information carried by the different wavelengthcomponents can be translated into a simultaneous or essentiallyinstantaneous “z-scanning” of the imaged z-depth range. This translationof the interference data into “z-scan” data can be carried out by animage generator 234 to generate and output an OCT image from the Fouriertransform of the sensed spectral components.

Some embodiments of the OCT camera 230 may use CCD (charge-coupleddevice) pixels as the sensors of the sensor array 232. Other embodimentscan achieve improved readout speeds by using CMOS sensors. In suchembodiments, the CMOS sensors can be read out in parallel. Further, inCMOS embodiments, it is possible to read out only sensors or pixels ofinterest, either selected prior to the imaging, or selected in real timebased on whether their content changed because of the imaging. Both ofthese aspects make CMOS pixels quite useful for speeding up theperformance of the OCT camera 230.

Using standard optical analysis, the critical imaging and performanceparameters of the SD-OCT system 200 can be characterized by itsarchitectural and design parameters as follows. Since the spectraldecomposer 231 directs the different wavelength components of thecombined beam 226 into slightly differing directions, the smaller andmore densely packed the individual sensors or pixels are (the smaller d′is), the narrower δλ wavelength/spectral ranges are resolved by the OCTcamera 230. The other quantity, determining δλ besides the pixeldensity, is the total range of wavelengths, i.e. the bandwidthW_(camera) of the OCT camera 230. In a simple arrangement, δλ isproportional to the bandwidth W_(camera) and inversely proportional tothe number of pixels in a row of the sensor array 232.

The important imaging z-depth range, or z-imaging range, Lmax isintimately related to δλ: the narrower the δλ wavelength ranges, thebroader the imaging range in the z direction because these twoquantities are connected by an inverting Fourier transform. Inparticular, the theoretical maximum imaging range is given by

$\begin{matrix}{{L\;\max} = {{\frac{1}{4}\left( \frac{\lambda_{0}^{2}}{\delta\;\lambda} \right)} = {\frac{1}{2}\frac{1}{Nf}}}} & (1)\end{matrix}$

Here, the value λ₀ refers to the average or central wavelength of thebroad-band light source 210 and Nf denotes the Nyquist frequency of theOCT camera 230. In reality, additional factors may limit the effectiveimaging range below this theoretical maximum, such as the signal tonoise ratio. Therefore, the effective imaging range is typically smallerthan this theoretical value Lmax.

One factor that can limit the imaging range further is the Rayleighrange R. R can be expressed using Δx, the resolution in the transverse xdirection, or “transverse resolution”, governed by the numericalaperture NA and the wavelength λ₀ of the light source 210. Specifically,Δx can be expressed as:

$\begin{matrix}{{\Delta\; x} = {\frac{4}{~\pi}\left( {\lambda_{0}\frac{f}{d}} \right)}} & (2)\end{matrix}$

where f is the focal length and d is the pupil of the objective 122,their ratio determining NA. Using Δx, the above discussed Rayleigh rangeR can be expressed as:

$\begin{matrix}{R = {\frac{\pi}{2}\left( \frac{\left( {\Delta\; x} \right)^{2}}{\lambda_{0}} \right)}} & (3)\end{matrix}$

The Rayleigh range R is often defined as the z directional distancebetween the focal depth and the depth where the beam's width is √{squareroot over (2)} times the width at the focal depth. Thus, R characterizesthe z-range within which the beam is narrow enough to enable highresolution imaging as limited by geometrical and wave optics. In thiscontext, Lmax can be thought of as characterizing the z-imaging range aslimited by the light source 210 and the resolution of the sensor array232. A system design principle often thought of as optimal, e.g. forGaussian beams, is to make these two z-ranges align with each other. Forexample, in some implementations, Lmax can be chosen to be essentiallyequal to 1-6 R:Lmax=1, . . . 6R  (4)

The same design principle can be stated through the concept of thewidely used “depth of focus”, which is often defined as twice theRayleigh range.

As shown by Eqs. (1)-(4), the z-imaging range depends on Lmax and R,which in turn depend on the system design parameters including λ₀, δλ,W_(camera), W_(source), f, d′, Nf, and d. Thus, for imaging systems forcataract surgery, the above system design parameters are to be chosensuch that the z-depth imaging range of the SD-OCT imaging system 200exceed 4 mm or 6 mm, such as to fall in the range of 4-20 mm or 6-10 mm,thus making the cataract surgical system 100 capable of assistingcataract surgeries by high resolution and sufficiently fast imaging.This design requirement is quite demanding and distinguishes cataractimaging systems from corneal or retinal imaging systems.

FIG. 7 illustrates an embodiment that can ensure not only a largez-depth imaging range, but a fast imaging time, allowing the SD-OCTimaging system 200 to provide feedback images in a timely and thusactionable manner, including operating at an essentially live videorate. As discussed above, a cataract surgical system 100 with the SD-OCTimaging system 200 can have its control system, including the OCT imageprocessor 201 and the laser controller 130 operate in essentially realtime, with the option of adjusting or modifying the surgical scanpatterns during the surgery according to the received feedback imaginginformation.

As described below in detail, embodiments of FIG. 7 are also configuredto scan the OCT imaging beam 308/222 particularly fast, usingprecomputed scan patterns. In some embodiments, these fast imagingrefresh rates of the SD-OCT imaging system 200 can be achieved byincluding a dedicated Input-Output board 260.

One function of the dedicated Input-Output board 260 is to addressproblems of some existing OCT imaging systems that do not have circuitryand a processor dedicated to imaging. In these systems, the processorthat is in charge of imaging can be forced or prompted to multitask andperform more than one function in an interleaved, parallel oroverlapping manner. To carry out these functions, the imaging processormay perform an “interrupt” by switching from e.g. the task of scanningthe imaging beam 222/308 to another task and back. Such interrupts,however short, can cause problems, since during the time when thescanning is stopped or frozen by the interrupt, the laser beam mayremain pointed to the same position. This scanning-freeze can disruptthe timing of the (x,y) scan, introducing an error and noise into thecoordinates of the imaged locations. This timing error in the outputtedscanning data can reach delays of 50, 100 or more microseconds: aphenomenon sometimes called jitter.

In addition, typically several other input/output agents communicate onthe same system bus on which the imaging processor is driving thescanning of the imaging beam, all demanding a fraction of the bus'scycle time. This shared nature of the channel allows it to support onlyslow data transfer rates, unfortunately. Further, to manage thesecompeting demands, a portion of the cycle of the system bus is typicallytaken up by control signals. Therefore, even if an OCT imaging system isdesigned to avoid the scanning-freeze by switching the imaging processorto outputting the scanning data to the scanning unit in a single-taskmode through a dedicated link, then the imaging processor will not beable to perform its other functions during this outputting step, such ascomputing the next scanning pattern. All these constraints slow down theperformance of such existing imaging systems considerably.

Implementations of the SD-OCT imaging system 200 can overcome thesedifficulties by employing the following efficient design. The scanningof the image beam 222 can be controlled by an imaging processor 250 anda dedicated Input-Output board 260. The imaging processor 250 cancompute scanning data such as the target scan pattern 302 and themodified scan pattern 312. These scanning data can include e.g. asequence of (x,y) coordinates where the OCT image beam 222 is to bedirected in the cataract target region. The imaging processor 250 cancompute the scanning data as well as perform its other functions inconnection to a storage medium that stores a computer code orinstruction set to facilitate these functions of the imaging processor250.

The dedicated Input-Output board 260 can include a local or dedicatedmemory controller 262, also referred to as a direct memory access (DMA)engine 262. The DMA engine/memory controller 262 can manage a transferof the computed scanning data, indirectly or directly, from the imagingprocessor 250 toward a data buffer 264. The data buffer 264, coupled tothe local memory controller 262 can store the scanning data and can beoperable to output the scanning data towards an output digital-analogconverter (output DAC) 266 at a high speed. The output DAC 266 can becoupled to the data buffer 264 to receive the scanning data, to convertselected outputted scanning data to analog scanning signals, and tooutput the scanning signals towards an OCT beam scanner 268 e.g. in ascanning data burst mode.

The image beam 222 can be scanned by the OCT beam scanner 268 through aseparate dedicated imaging optic, or partially through the guiding optic120 of the surgical beam. In either of these implementations, the imagebeam 222 can be coupled into the eye through the objective 122 and thecorresponding docking unit or patient interface (PI) 124. In otherembodiments, the image beam 222 can be guided into the eye 1 through airwithout the docking unit 124 being docked to the eye 1.

The output of the scanning data by the output DAC 266 can besynchronized by an imaging sync 242 to the operation of the OCT camera230, so that the OCT camera 230 can take the OCT images synchronouslywith the scanning operations. The synchronously taken OCT images can beoutputted to the OCT image processor 201 that can perform any one of thelarge number of image processing tasks described up to now. Finally, thegenerated and processed images can be displayed by an OCT image display270. In some embodiments, the imaging processor 250 and the OCT imageprocessor 201 can be integrated partially or completely.

FIG. 8 illustrates an implementation of the dedicated Input-Output board260 in some more detail. The imaging processor 250 can be coupled to abus 252, such as a PCI bus 252. The system can also include a processormemory 254. The imaging processor 250 can compute the scan patterns andthen output the computed scan patterns through the shared PCI bus 252 tothe processor memory 254. After the imaging processor 250 generated thescan patterns but before the commencement of the actual scan operation,the dedicated DMA engine 262 can transfer the scanning data from theprocessor memory 254 to the data buffer 264. The data buffer 264 can bea first-in-first-out (FIFO) memory 264. The FIFO data buffer 264 canstore the scan pattern or scanning data and output the stored scanningdata to the output DAC 266 when prompted by the dedicated DMA engine262. The output DAC 266 can convert the scanning data into analogscanning signals and output them to an x galvo beam scanner 268 x and ay galvo beam scanner 268 y of the OCT beam scanner 268 that control xand y galvo mirrors, or redirector elements, to scan the OCT image beam222/308 according to the target scan pattern 302 and the modified scanpattern 312, coded in the scanning data. Some implementations may havean integrated (x,y) galvo-controller 268 xy that controls a single galvomirror capable of rotating around both the x and y axes. The output DAC266 can also drive the image sync 242 to synchronize the taking of theOCT images with the scanning operations.

In some implementations, the imaging processor 250 can output thescanning data to the dedicated Input-Output board 260 through adedicated memory bus or through a local bus instead of the shared PCIbus 252. In other implementations, there can be even a direct connectionbetween the imaging processor 250 and the DMA engine 262.

This design is efficient at least for the following reasons. (1) Thescanning data or scan patterns are pre-computed by the imaging processor250, thus no time consuming real-time scanning data computation isinvolved. (2) The imaging processor 250 is not tasked with outputtingthe scanning data in real time, as the pre-computed scanning data arestored in the dedicated data buffer 264. This design can reduceinterrupts, freezes and jitters below 50, 40, or even 20 microseconds,caused by the imaging processor 250 multitasking. (3) The transfer ofthe scanning data will not be interrupted by the bus 252 being shared byother agents, neither will it be slowed down by the typically slowtransfer rates of the shared PCI bus 252. (4) The data buffer 264 isdedicated to the task of scanning, so the output of the scanning datacan be performed in a fast transfer mode, such as a burst mode, furtheraccelerating the scanning speed.

In addition, since the dedicated Input-Output board 260 drives theoutputting of the scanning data essentially autonomously, the imagingprocessor 250 is free to perform other functions in parallel with thescanning data output, such as generating the modified scan pattern 312.

In some implementations, the speed of the output by the output DAC 266can be so fast that an operating speed of the SD-OCT imaging system 200can be limited by an integration time of the OCT camera 230 instead ofthe speed of the scanning electronics. In some of these implementations,the output DAC 266 can output the scanning signals at a rate within oneof the following ranges: 1 Hz-1 MHz, 100 Hz-1 MHz, or 1 kHz-100 kHz.

FIG. 9 illustrates that some cataract surgical laser systems 100 caninclude another type of imaging system: a Swept-Source-OCT (SS-OCT)imaging system 280. The SS-OCT imaging system 280 can include a sweptwavelength light source 282 that emits a coherent image beam with anarrower bandwidth W_(source), than the SD-OCT light source 210. Byadvanced modulation techniques the SS-OCT light source 282 can vary thewavelength of the emitted image beam 222, “sweeping” the wavelength λacross the bandwidth W_(source) in time. The SS-OCT imaging system 280can employ a beam guidance system 211 that is analogous to that of theSD-OCT imaging system 200. In particular, the 1^(st) beam splitter 212can again create the combined beam 226 that carries the imaginginformation associated with different wavelengths.

As a difference from the spectrometer-based imaging systems, the SS-OCTimaging system 280 separates the different wavelengths or spectralcomponents in time, whereas the SD-OCT systems 200 separate them inspace. The different wavelength components, carrying image datacorresponding to different z-depths, are separated into a time sequenceas the wavelength λ □ is swept by the SS-OCT light source 282.Therefore, the OCT camera 230 of the SS-OCT systems 280 is different aswell.

In some implementations, it consists of a single detector 284 that candetect and resolve the combined beam 226 in very short time intervals.In some embodiments, the detector 284 can be an avalanche photo-diode ora photomultiplier tube. The detector 284 can be capable of transferringor dumping the detected signals, corresponding to different wavelengthsor spectral components, to a set of data binners 286. Some embodimentsof the SS-OCT imaging system 280 are analogous to the SB-OCT imagingsystems because both of them generate the images via spectraldecomposition. The spectral components of the SS-OCT image can beassembled into the OCT image similarly as in the SB-OCT systems: a FastFourier Transformer 288 can perform a Fourier transformation of thecontents of the data binners 286 to assist the image generator 234 togenerate the OCT image. The FFT unit 288 can be analogous to the FFTunit 233 in the SD-OCT imaging system 200.

According to the above description, the SS-OCT imaging systems 280 havefeatures similar to the TD-OCT imaging systems as at one phase theimaging data is captured sequentially, not in parallel. However, unlikein TD-OCT systems, the different z-depth imaging data are captured withdifferent spectral components of the combined beam 226, necessitatingthe performing of the Fourier transformation by the FFT unit 288. Inthis sense, the SS-OCT imaging systems 280 are related to the SD-OCTimaging systems 200 that manifestly work with different spectralcomponents. SS-OCT systems are close to the SD-OCT systems in one moresense: they sweep the wavelength of the image beam of the sweptwavelength light source 282 without moving mechanical parts such as thereference mirror 216. Finally, as the sweeping of the wavelength of theswept wavelength light source 282 can be performed with a speed muchabove the scanning speed of TD-OCT system as no moving parts areinvolved in the sweeping, SS-OCT systems 280 can image at speeds muchfaster than TD-OCT systems, albeit below the imaging speeds of theSD-OCT systems. Therefore, implementations of the SS-OCT imaging system280 can also be able to generate their images at live refresh rates withacceptable resolution, providing a very useful functionality andactionable feedback for the cataract surgical system 100.

FIG. 10 illustrates an ophthalmic surgical method 500 to operate thecataract surgical system 100. The method 500 can include: a generatingan electronic representation of a target scan pattern 302 for the lens 5of the eye 1 by the laser controller 130 (510); a generating andscanning a surgical laser beam 304 in the lens 5 of the eye according tothe target scan pattern 302 by the cataract surgical laser system 100,creating a cut 306 in the lens (520); a generating an image of a portionof the eye and the cut 306 with a Spectral Domain Optical CoherenceTomographic imaging system 200 after the scanning of the laser beamstarted (530); a generating an electronic representation of a modifiedscan pattern 312 by the laser controller 130 in relation to thegenerated image (540); and a generating and scanning the surgical laserbeam 304 in the lens 5 of the eye according to the modified scan pattern312 by the cataract surgical laser system 100, creating a modified cut314 (550).

In some implementations, the generating an electronic representation ofa modified scan pattern 540 can include receiving a modification inputfrom a system operator in response to the generated image of the portionof the eye.

In other implementations, the generating an electronic representation ofa modified scan pattern 540 can include analyzing the generated image bythe OCT image processor 201; determining a deviation of the cut 306relative to the target scan pattern 302; and generating a control signalby the OCT image processor 201 for the laser controller 130 to generatethe modified scan pattern 312.

FIG. 11 illustrates a related method of cataract surgery 600. The method600 can include controlling a scanning of the surgical laser beam 304 inthe lens 5 of the eye 1 by the laser controller 130 (610); generatingimages of a portion of the lens 5 at a rate of at least 5 frames persecond by the Spectral Domain Optical Coherence Tomographic imagingsystem 200 (620); and modifying the scanning of the surgical laser beam304 by the laser controller 130 in response to an analysis of thegenerated images (630).

While this specification contains many specifics, these should not beconstrued as limitations on the scope of the invention or of what can beclaimed, but rather as descriptions of features specific to particularembodiments. Certain features that are described in this specificationin the context of separate embodiments can also be implemented incombination in a single embodiment. Conversely, various features thatare described in the context of a single embodiment can also beimplemented in multiple embodiments separately or in any suitablesubcombination. Moreover, although features can be described above asacting in certain combinations and even initially claimed as such, oneor more features from a claimed combination can in some cases be excisedfrom the combination, and the claimed combination can be directed to asubcombination or variation of a subcombination.

The invention claimed is:
 1. An ophthalmic surgical system, comprising:a laser source, configured to generate a first set of laser pulses; alaser controller comprising a processor configured to: generate anelectronic representation of a first scan pattern; and cause a laserscanner to scan the first set of laser pulses according to a portion ofthe first scan pattern to create a first photo-disrupted region in atarget region of an eye; and an Optical Coherence Tomographic (OCT)imaging system, configured to generate an image that includes a portionof the first photo-disrupted region with an image resolution in therange of 0.5-10 million image points per image and a frame-rate in therange of 20-500 frames/sec; wherein the processor of the lasercontroller is further configured to: generate an electronicrepresentation of a modified scan pattern based on the image generatedby the OCT imaging system; and cause the guiding optic to scan a secondset of laser pulses according the modified scan pattern to create asecond photo-disrupted region.
 2. The system of claim 1, wherein the OCTimaging system comprises a Swept-Source OCT (SS-OCT) imaging system. 3.The system of claim 2, wherein the SS-OCT imaging system is configuredto generate the image that includes a portion of the firstphoto-disrupted region with an image resolution in the range of 0.5-2million image points per image and a frame-rate in the range of 20-200frames/sec.
 4. The system of claim 2, wherein the SS-OCT imaging systemis configured to generate the image that includes a portion of the firstphoto-disrupted region with an image resolution in the range of 2-10million image points per image and a frame-rate in the range of 25-500frames/sec.
 5. The system of claim 2, wherein the SS-OCT imaging systemis configured to generate the image that includes a portion of the firstphoto-disrupted region with a resolution in the range of 2,000-5,000A-scans per B-scan.
 6. The system of claim 2, wherein the SS-OCT imagingsystem is configured to generate the image that includes a portion ofthe first photo-disrupted region with an A-scan acquisition rate of30-300 kHz.
 7. The system of claim 2, wherein the SS-OCT imaging systemis configured to generate the image that includes a portion of the firstphoto-disrupted region with an A-scan acquisition rate of 100-1,000 kHz.8. The system of claim 2, wherein: the target region comprises ananterior capsular layer; and the first scan pattern comprises a set ofpoints on a cylinder to form at least one of a circular capsulotomy, ananterior capsulotomy, and a curvilinear capsulotomy.
 9. The system ofclaim 2, wherein: the target region comprises a portion of a crystallinelens of the eye; and the first scan pattern comprises a set of points onat least one of radial chop planes, cylinders, a spiral pattern and amesh pattern to induce at least one of a chop, a photo-disruption and alysis of the lens.
 10. The system of claim 2, wherein the SS-OCT imagingsystem is configured to have a z-imaging range greater than 4 mm. 11.The system of claim 2, wherein the SS-OCT imaging system is configuredto have a z-imaging range greater than 6 mm.
 12. The system of claim 2,wherein the SS-OCT imaging system is configured to generate the image inan imaging time less than 0.1 sec.
 13. The system of claim 2, whereinthe SS-OCT imaging system comprises: a swept wavelength light source togenerate a swept-wavelength beam; a beam guidance system, configured to:split the swept-wavelength beam into an image beam and a reference beam,guide the image beam to the eye and to guide a returned image beam fromthe eye, guide the reference beam to a reference mirror and to guide areturned reference beam from the reference mirror, and combine thereturned image beam and the returned reference beam into a combinedbeam; and an OCT camera configured to receive the combined beam,comprising: a detector to detect the combined beam; a data binner todetect the combined beam as a time sequence of data; aFast-Fourier-Transform-system to Fourier transform the detected timesequence of data; and an image generator to generate an image from theFourier transform.
 14. The system of claim 2, wherein the processor ofthe laser controller is configured to generate the electronicrepresentation of the modified scan pattern according to a modificationinput received from a system operator in response to the system operatorhaving analyzed the image of the portion of the first photo-disruptedregion.
 15. The system of claim 1, wherein the SS-OCT imaging system isconfigured to output calibration marks for display to a system operatorfor calibrating a location of the first photo-disrupted region relativeto the first scan pattern.